(12) INTERNATIONAL APPLICATION PUBLISHED UNDER THE PATENT COOPERATION TREATY (PCT)
(19) World Intellectual Property Organization
International Bureau
(43) International Publication Date
22 November 2001 (22.11.2001)
PCT
(10) International Publication Number
WO 01/88524 Al
(51) International Patent Classification 7 : COIN 27/26
(21) Internationa! Application Number: PCT/US01/15019
(22) International Filing Date: 10 May 2001 (10.05.2001)
(25) Filing Language: English
(26) Publication Language: English
(30) Priority Data:
60/203,762
12 May 2000 (12.05.2000) US
(71) Applicant: THERASENSE, INC [US/US]; 1360 South
Loop Road, Alameda, CA 94502 (US).
(72) Inventors: HELLER, Adam; 4711 Spicewood Springs
Road, Apt. 271, Austin, TX 78759 (US). CHEN, Ting;
2501 Lake Austin Boulevard #D102, Austin, TX 78703
(US). FRIEDMAN, Keith, A.; 500 East Riverside Drive
#256, Austin, TX 78704 (US).
y
(74) Agent: BRUESS, Steven, C; Merchant & Gould P.C.,
P.O. Box 2903, Minneapolis, MN 55402-0903 (US).
(81) Designated States (national): AE, AG, AL, AM, AT, AT
(utility model), AU, AZ, BA, BB, BG, BR, BY, BZ, CA,
CH, CN, CO, CR, CU, CZ, CZ (utility model), DE, DE
(utility model), DK, DK (utility model), DM, DZ, EC, EE,
EE (utility model), ES, Fl, FI (utility model), GB, GD, GE,
GH, GM, HR, HU, ID, XL, IN, IS, JP, KE, KG, KP, KR, KZ,
LC, LK, LR, LS, LT, LU, LV, MA, MD, MG, MK, MN,
MW, MX, MZ, NO, NZ, PL, PT, RO, RU, SD, SE, SG, SI,
SK, SK (utility model), SL t TJ, TM, TR, TT, TZ, UA, UG,
UZ, VN, YU, ZA, ZW.
(84) Designated States (regional): ARIPO patent (GH, GM,
KE, LS, MW, MZ, SD, SL, SZ, TZ, UG, ZW), Eurasian
patent (AM, AZ, BY, KG, KZ, MD, RU, TJ, TM), European
patent (AT, BE, CH, CY, DE, DK, ES, FI, FR, GB, GR, IE,
IT, LU, MC, NL, PT, SE, TR) f OAPI patent (BF, BJ, CF,
CG, CI, CM, G A, GN, GW, ML, MR, NE, SN, TD, TG).
Published:
— with international search report
For two-letter codes and other abbreviations, refer to the "Guid-
ance Notes on Codes and Abbreviations" appearing at the begin-
ning of each regular issue of the PCT Gazette.
(54) Title; ELECTRODES WITH MULTILAYER MEMBRANES AND METHODS QF USING AND MAKING THE ELEC-
TRODES
100
es
IT)
00
00
102 104 /
/ / / 106 108 110
(57) Abstract: A sensor including a sensing layer (106) disposed over an electrode (102) or an optode and a layer-by-layer assembled
mass transport limiting membrane (110) disposed over the sensing layer. The membrane includes at least one layer of a polyanionic
or polycationic material. The assembled layer of the membrane are typically disposed in an alternating manner. The sensor also
^ optionally includes a biocompatible membrane (110).
WO 01/88524 PCT/US01/15019
1
ELECTRODES WITH MULTILAYER MEMBRANES AND METHODS OF
USING AND MAKING THE ELECTRODES
This application is being filed as a PCT International Patent Application in
5 the name of TheraSense, Inc., a U.S. national corporation and resident, on 10 May
2001, designating all countries except the U.S.
Field of the Invention
This invention relates to sensors and sensor components that have multilayer
membranes and methods of making and using the sensors and sensor components. In
10 addition, the invention relates to enzyme electrodes and optodes with multi-layer
analyte-flux limiting membranes and methods of making and using the optodes and
the electrodes.
Background of the Invention
Miniature biosensors utilizing enzyme-containing optodes and electrodes for
15 monitoring biochemicals often include mass-transport controlling membranes. The
membranes can affect some or all of the characteristics of the optodes or electrodes,
including their sensitivity, size, apparent stability, dynamic range and selectivity.
Micro-membranes for use with miniature biosensors typically cannot be easily cut to
size from pre-formed membranes and if cut by a precision tool, such as a laser or an
20 electron beam, their placement on and attachment to the surface of an electrode or
optode can be difficult.
The reproducible casting of micro-membranes can also be difficult. For cast
micro-membranes, the pore sizes and their distribution are typically determined by
the relative rates of nucleation and mass-transport during generation of the
25 membrane by phase separation as a result of solvent evaporation. The outcome of
the simultaneously occurring nucleation and the mass transport processes depends
on the evolution (meaning the time-dependence during the solvent evaporation) of
the viscosity, the concentrations of the solvent and the non-solvent, and the
membrane's leached phase and its residual phase. These are affected by time-
30 dependent temperature gradients and by the time-dependent gradient of the partial
WO 01/88524 PCT/US01/15019
2
pressure of the evaporating solvent over a droplet, the dimensions of which shrink
and are a function of the time-dependent contact angle with the wetted surface.
Summary of the Invention
Generally, the present invention relates to electrodes and optodes having
5 membranes to reduce analyte flux or reduce interferent flux or both. One
embodiment is a sensor that includes a sensing layer disposed on a substrate and a
multilayer flux- limi ting membrane disposed over the sensing layer. The membrane
includes a first layer disposed on and bound to the sensing layer and one or more
additional layers disposed on and bound to the preceding layers of the membrane.
1 0 The substrate can have a conductive material upon which the sensing layer is
disposed to form an electrode or an optical material, such as an optical fiber, upon
which the sensing layer is disposed to form an optode. As an example, the
membrane includes at least two layers; one of which is a polycationic layer or a
polyanionic layer. Optionally, the membrane includes at least one layer that has
15 functional groups that can capture transition metal ions.
Another embodiment is a method of making a sensor. A sensing layer is
disposed on a substrate. A first membrane layer is disposed on and binds to the
sensing layer. One or more subsequent membrane layers are disposed over the first
membrane layer, each of the subsequent membrane layers binding to the
20 immediately preceding membrane layer. For example, the membrane layers can be
formed by chemisorption or reactive adsorption.
The above summary of the present invention is not intended to describe each
disclosed embodiment or every implementation of the present invention. The
Figures and the detailed description which follow more particularly exemplify some
2 5 but not all of these embodiments .
Brief Description of the Drawings
The invention may be more completely understood in consideration of the
following detailed description of various embodiments of the invention in
connection with the accompanying drawings, in which:
WO 01/88524 PCT/US01/15019
3
Figure 1 is a graph illustrating the dependence of the sensitivity of a glucose
sensing electrode on the number of PAc/PAm/PAc/PAm/PAc/PVPEA sextets:
(dashed line) none; (squares) one; (triangles) two; (circles) three;
Figure 2 is a graph illustrating dependence of the apparent stability of a
5 glucose sensing electrode on the number of PAc/PAm/PAc/PAm/PAc/PVPEA
sextets: (a) none; (b) two; (c) three;
Figure 3 is a graph illustrating current increments upon raising the
concentrations sequentially by 5 mM glucose, 5 mM glucose, 0.1 mM ascorbate, 0.2
mM acetaminophen and 0.5 mM urate;
10 Figure 4 is a graph illustrating changes in the current upon raising the
glucose concentration in 5 mM increments, then adding Fe 2+ in an amount that in the
absence of precipitation of iron phosphate would have raised the concentration of the
cation to 0.1 mM with (a) ten PAc/PAm bilayers or (b) three
PAc/PAm/PAc/PAm/PAc/PVPEA sextets;
15 Figure 5 is a graph depicting in vivo experiments in which the glucose
concentration was tracked by sensors with three PAc/PAm/PAc/PAm/PAc/PVPEA
sextets, (a) ( — ) output of the sensor implanted in the jugular vein; (...) output of the
sensor implanted in the intrascapular subcutaneous tissue; (triangles) concentrations
of glucose in blood samples withdrawn from the contralateral jugular vein measured
20 with a YSI glucose analyzer, (b) results of linear regression analysis of the data from
the two sensors in figure 5(a): in the jugular vein (closed circles) and in the
intrascapular subcutaneous tissue (open circles);
Figure 6 is a Clarke-type clinical error grid diagram of all data points
assuming that the measured glucose concentration does not lag behind the blood
25 glucose concentration at any time;
Figure 7 is a Clarke-type clinical error grid diagram of all data points
assuming that the sensor-measured glucose concentration lags by 3 minutes behind
the blood glucose concentration in the period of rise after intravenous injection of a
bolus of glucose and lags by 9 minutes during the period of decline after intravenous
30 injection of insulin; and
Figure 8 is a schematic cross-sectional view of one embodiment of an
electrode, according to the invention.
WO 01/88524 PCT/US01/15019
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While the invention is amenable to various modifications and alternative
forms, specifics thereof have been shown by way of example in the drawings and
will be described in detail. It should be understood, however, that the intention is
not to limit the invention to the particular embodiments described. On the contrary,
5 the intention is to cover all modifications, equivalents, and alternatives falling within
the spirit and scope of the invention.
Detailed Description of the Preferred Embodiment
The present invention is believed to be applicable to sensors, particularly
miniature sensors, and methods of making and using the sensors. In particular, the
10 present invention is directed to membranes for electrodes and optodes for use in
sensors and methods of making and using the electrodes and optodes. While the
present invention is not so limited, an appreciation of various aspects of the
invention will be gained through a discussion of the examples provided below.
A sensor includes an electrode or optode. Each electrode or optode includes
1 5 a substrate. For an electrode, the substrate includes a conductive material that is
typically either formed on a non-conductive support (e.g., a polymeric film) or as a
wire, rod, plate, or other object. A sensing layer (e.g., transducing layer) is provided
over at least a portion of the conductive material to transduce the flux of a chemical,
usually a biochemical, termed the analyte to an electrical signal, such as a current
20 flowing through the electrode. An optode can include, for example, an optical fiber
(or other optical substrate) at the tip of which a layer of a biologically active
macromolecule is provided (preferably, immobilized). The sensing layer on the
optode converts the change in the concentration of the analyte to a change in photon
flux.
25 The electrode or optode is covered with a membrane. The purpose of the
membrane is to reduce analyte-flux or reduce or prevent interferent flux to the
electrode or a combination of these features. The membrane contains at least two
layers, and, typically, at least six layers. At least one of the layers includes a layer
containing a polyanion or a polycation and the polyanionic or polycationic layer is
30 disposed over the sensing layer. When operating to reduce flux of an analyte,
interferent, or other material, the membrane typically provides a thin zone in which
WO 01/88524 PCT/US01/15019
5
the solubility of a solute (e.g., analyte or interferent) is lower by at least an order of
magnitude than it is in the sample solution. A membrane having such a zone is
referred to as "mass transport limiting". The cross sectional area of the sensor
through which mass transport is limited is defined as the "active area". The active
5 area of the sensors in at least some embodiments of this invention is 0. 1 cm 2 or
smaller. For example, sensors can be formed with active areas between 10" 2 cm 2 and
10* 8 cm 2 . Membranes having mass transporting areas (e.g., "active areas") between
10" 3 cm 2 and 10~ s cm 2 are particularly useful for biosensing applications.
The membrane can be formed, for example, by sequential chemisorption of
1 0 layers. Preferably, each particular layer, other than the terminal top layer, binds both
the preceding and succeeding layers of the membrane. Generally, two consecutive
layers are not identical; however, layers made of the same material but differently
(for example, oppositely) oriented can be disposed next to each other. The layers
typically form an array of bonds as a result of ionic, hydrophobic, coordinative,
1 5 covalent, van der Waals, or hydrogen bonding interactions between the materials of
the two layers. For example, the membrane can be formed of alternating
polyanionic and polycationic layers.
In one embodiment, a micro-membrane is formed in-situ on a miniature
enzyme electrode by disposing (e.g., chemisorbing or otherwise depositing) a
20 polyanionic material on a polycationic surface of the electrode, rinsing, disposing a
polycationic on the polyanionic material, rinsing, and repeating the cycle a desired
number of times. It will be understood that a similar procedure can be used with a
polyanionic surface of the electrode by first depositing a polycationic material
followed by a polyanionic material and repeating the cycle a desired number of
25 times. Other orders of layers can also be used.
In particular, the membranes are useful in those biosensors that function by
chemically converting (e.g., reacting) a chemical or biochemical. The chemical or
biochemical can be, for example, an analyte that is being assayed, a product formed
by reaction of the analyte, a co-reactant of the analyte, a product or reactant of a
30 reaction that is catalyzed or inhibited by the presence of the analyte, or a constituent
whose attachment to an optode or electrode is accelerated or inhibited by the
presence of the analyte.
WO 01/88524 PCT/US01/15019
6
As an example, membranes of the invention can be formed on, and are
evaluated in the Examples below for, miniature glucose oxidizing or
electrooxidizing electrodes. For these electrodes, glucose is typically converted to
gluconolactone in the first step of a detection reaction. Suitable miniature glucose
5 electrodes include, for example, those disclosed in U.S. Patents Nos. 5,262,305;
5,262,035; 5,264,104; 5,593,852; and 5,665,222 and U.S. Patent Applications Nos.
08/795,767; 09/034,372; 09/295,962; and 09/434,026, all of which are incorporated
herein by reference. It will be appreciated that electrodes for the detection of
analytes other than glucose can also benefit from the use of the membranes
10 described herein. Glucose electrodes are illustrated herein as an application
example.
Glycemia has been monitored amperometrically in the subcutaneous
interstitial fluid with miniature electrodes for some time. The electrode reactions
applied in such monitoring include (a) the mediated electrooxidation of glucose to
1 5 gluconolactone at electrodes coated with a redox mediator (e.g., a redox polymer
which electrically "wires" the reaction centers of glucose oxidase to an electrode)
(Equations la and lb, below) or, alternatively, (b) the glucose oxidase catalyzed
reaction of glucose with 0 2 , to produce gluconolactone and H 2 0 2 (Equation 2a,
below), followed either by electro-oxidation of the H 2 0 2 (Equation 2b, below), or by
20 monitoring the change in the 0 2 partial pressure or concentration.
glucose + 2 bound mediator(ox) -> gluconolactone + 2 bound mediator(red) + 2H +
2 bound mediator (red) -» 2 bound mediator (ox) + 2e~
glucose + 0 2 -> gluconolactone + H 2 0 2
25 H 2 0 2 -* 0 2 + 2H* + 2e-
The current at the electrode generally scales with the flux of glucose to the electrode.
The flux, and therefore also the current, typically increases linearly with the glucose
concentration as long as the entire glucose flux at the electrode is consumed in the
3 0 electrode reaction.
Figure 8 illustrates one example of a suitable electrode 100. The electrode
includes a conductive material 102 within an insulating sleeve 104. Disposed on the
(la)
(lb)
(2a)
(2b)
WO 01/88524 PCT/US01/15019
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conductive material 102 are a sensing layer 106, a membrane 108, and a
biocompatible layer 110. Examples of similar electrode configurations, together
with examples of some suitable sensing layers and biocompatible layers are
provided in U.S. Patent No. 5,593,852. Other suitable sensing layers and
5 biocompatible layers are described, for example, in U.S. Patent No. 5,665,222 and
U.S. Patent Applications Nos. 08/795,767; 09/034,372; 09/295,962; and 09/434,026,
all of which are incorporated by reference.
The conductive material 102 can be, for example, carbon, a metal, or a
conductive compound or polymer. The insulating sleeve 104 is typically formed
10 from an insulating compound or polymer. In one method of manufacture, a tip of a
metal wire (e.g., a gold wire) is etched to form a recess within the insulating sleeve
where the sensing layer 106, membrane 108, and biocompatible layer 110 are
deposited.
The sensing layer 106 typically provides a mechanism for transducing an
15 analyte flux (or the flux of a product molecule formed or consumed by a reaction of
the analyte) to an electrical signal. The sensing layer can contain, for example, a
redox mediator to facilitate the indirect or direct transfer of electrons between the
conductive material and the analyte. One type of redox mediator is a transition
metal complex or compound (e.g., an osmium, ruthenium, or iron complex or
20 compound). The redox mediator can be a monomeric redox compound or complex,
but is preferably in a non-leachable form, such as a redox polymer. The redox
polymer, has a polymeric backbone with multiple redox centers. Examples of
suitable redox mediators are disclosed in, for example, U.S. Patents Nos. 5,262,305;
5,262,035; 5,264,104; 5,593,852; and 5,665,222 and U.S. Patent Applications Nos.
25 08/795,767; 09/034,372; 09/295,962; and 09/434,026, all of which are incorporated
by reference.
The sensing layer 106 can also include a second electron transfer agent, such
as an enzyme. The second electron transfer agent can catalyze the electrochemical
oxidation or reduction of the analyte. As an example, suitable second electron
30 transfer agents for glucose include glucose oxidase or glucose dehydrogenase; for
lactate, lactate dehydrogenase; and for hydrogen peroxide, peroxidase.
WO 01/88524 PCT/US01/15019.
The biocompatible layer 110 prevents the penetration of large biomolecules
into the electrodes. This can be accomplished by using a biocompatible layer 1 10
having a pore size that is smaller than the biomolecules that are to be excluded.
Such biomolecules can foul the conductive material 102 or the sensing layer 106
5 thereby reducing the effectiveness of the electrode 100 and altering the expected
signal amplitude for a given analyte concentration. The biocompatible layer 1 10
may also prevent protein adhesion to the electrode 100, formation of blood clots, and
other undesirable interactions between the electrode 100 and body. A preferred
biocompatible coating is a hydrogel, which contains at least 20 wt.% fluid when in
10 equilibrium with the analyte-containing fluid. Examples of suitable hydrogels are
described in U.S. Patent No. 5,593,852, incorporated herein by reference, and
include crosslinked polyethylene oxides, such as polyethylene oxide tetraacrylate.
The preferred membrane 108 includes at least three, and typically at least six,
twelve, or eighteen layers. One example of a membrane 108 is formed using
15 alternating polycationic and polyanionic layers. Typically, these layers are formed
using polymers. Suitable polycationic polymers include, for example,
polyallylamine hydrochloride (PAm), po!y(4-vinyIpyridine) quaternized by reacting
about one third to one tenth of the pyridine nitrogens with 2-bromoethylamine
(PVPEA), polyethylene imine, and polystyrene modified with quaternary
20 ammonium functions. Suitable polyanionic polymers include, for example,
poly(acrylic acid) (PAc), poly(methacrylic acid), partially sulfonated polystyrene,
polystyrene modified with functions having carboxylate anions, and DNA
(deoxyribonucleic acid) or RNA (ribonucleic acid) strands, fragments or oligomers.
The membrane 108 can serve one or more functions including, for example, a)
25 limiting glucose flux or b) reducing or eliminating the flux of interferents to the
electrode. Glucose sensors will be described as an example, but sensors can be
formed for other analytes using the same principles.
In glucose sensing optodes or electrodes , a glucose flux-limiting membrane
can enhance one or more properties of the optode or the electrode including, for
30 example, expanding the dynamic range, enhancing the apparent stability and
improving the selectivity for glucose which may enable one-point calibration of an
implanted electrode.
WO 01/88524 PCT/US01/15019
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Reducing the flux of glucose (or any other analyte) using a membrane can
expand the dynamic range of a sensor. The upper limit of the dynamic range of the
sensor is that glucose concentration where the entire influx of the analyte is still
consumed. When this limit is exceeded, the rate of glucose-conversion is slower than
5 the influx and the sensor "saturates": the current no longer increases when the
glucose concentration is raised. The kinetic limit, represented by the current density
at the glucose concentration above which it no longer increases, is an intrinsic
property of the electrode. This current density typically scales linearly with the rate
of the slower of reactions 1(a) or 1(b) for a mediator-comprising glucose oxidase
1 0 electrode or the slower of reactions 2(a) or 2(b) in an 0 2 - utilizing electrode.
Because the membrane does not affect the intrinsic rate of the slowest step, its
insertion between the assayed fluid and the electrode expands the upper limit of the
dynamic range.
Typically, the sensor's apparent stability, which is the stability perceived by
15 its user, can also be improved upon insertion of the glucose flux limiting membrane.
The user remains unaware of the deterioration of the electrode's chemistry as long as
the kinetics of the slowest step remains fast enough to convert all the influx of
glucose to a current. Furthermore, the more blocking the membrane is, the better the
apparent stability will typically be. The increased apparent stability and the upward
20 extension of the dynamic range are gained at the cost of reduced specific sensitivity,
defined as the current per unit area at unit glucose concentration. The specific
sensitivity decreases because less glucose reaches the reactive zone on the electrode.
When the specific sensitivity decreases, a larger glucose-transporting area may be
needed for the sensor's current to reach an easily measured value.
25 The membrane thus defines, at least in part, (a) the lowest analyte
concentration where the current is large enough to be easily measured and the
highest analyte concentration above which the electrode's current no longer
increases with the concentration of the analyte; (b) the extent of the feasible
miniaturization of the sensor; and (c) the apparent stability of the sensor. Fitting of
30 these characteristics to those sought in the intended application requires tailoring of
the flux-controlling membrane. While the characteristics of the earlier solvent-cast
membranes depended on the spreading of the casting solution, and on the gradients
WO 01/88524 PCT/US01/15019
10
of temperature and of solvent partial pressure near the membrane being formed; the
characteristics of the membranes formed according to the invention typically do not
depend on these factors. As a result, the simultaneous tailoring of their many desired "
characteristics is typically easier.
5 In at least some embodiments, the membrane also reduces or eliminates the
. flux of interferents to the conductive material of the electrode. The bias resulting
from the presence of an interferent that is rapidly electrooxidized at the applied
potential is defined by its flux to the electrode surface. In the absence of a
membrane, the flux is controlled by diffusion in the solution and is defined by the
10 diffusivity and the concentration of the interferent. When a membrane is inserted
between the solution and the electrode, the % bias is defined by the ratio of the
products of the solution concentrations and permeabilities of the interferent and
glucose.
%bias= Cj > stDj > s xl00% (solution)
1 5 %bias = Ci > s ' Pjm x 1 00% (membrane)
Cg,s ' Pgm
Q
P = K IM D IM l l M= X D IM ' l M
where C is the concentration of the interferent (I) or glucose (G) in the solution (S)
or membrane (M), D represents the appropriate diffusivity, P represents the partition
coefficient, and 1 M represents the thickness of the membrane. Because the
20 permeability is a product of the concentration of the diffusing species in the
membrane and of its diffusivity in the membrane, the % bias increases linearly with
the ratio of the partition coefficients of the interferent and of glucose between the
membrane and the solution. When only monovalent anionic species (CI", ascorbate,
urate) are present, the concentration of anions in the polycationic POs-EA membrane
25 of the sensing layer equals the concentration of its cationic charges, which is about
an order of magnitude higher than the concentration of anions in the solution. As a
result, the sensing layer is permselective for anions over glucose and in the absence
of a neutral or polyanionic glucose-flux-controlling membrane (e.g. Nafion™) the
flux and electro-oxidation current of 0. 1 mM ascorbate could equal or exceed 10%
WO 01/88524 PCT/US01/15019
11
of that of 10 mM glucose. The membranes described herein are not typically
permselective for anions. As a result, ascorbate or urate is not preferentially
electrooxidized. As seen in Figure 3, at 10 mM glucose concentration the combined
% bias resulting of the presence of 0.1 mM ascorbate, 0.2 mM acetaminophen and
5 0.5 mM urate is less than 5%, even though the sensor is poised at 450 mV
(Ag/AgCl), a potential where acetaminophen is not rapidly electrooxidized, but
ascorbate and urate are,
Transition metal ions also influence the sensor readings. Transition metal
ions reduce the intrinsic kinetic capacity of the enzyme layer to electrooxidize
10 glucose and thereby severely reduce the dynamic range and the sensitivity. (See
Figure 4) The loss is attributed to both inhibition of the enzyme and a reduction in
the surface density of electroactive redox centers caused by excessive crosslinking
through coordination of pyridine rings of neighboring redox centers. In the
Examples, a transition metal ion capturing PVPEA/PAc bilayer is included in the
15 membrane. The addition of the PVPEA layer provides an alternative site for capture
of transition metal ions by providing pyridine functionalities that can complex with
the transition metal ions. It will be recognized that other materials can be used in
place of PVPEA. Typically, the replacement materials will include functionalities
that can form complexes with the transition metal ions. Because the PVPEA layer is
20 already highly crosslinked, the incremental crosslinking by coordination of the
transition metal ions does not change excessively the permeability of the micro-
membrane to glucose.
As an example, described in the Examples section below, membranes were
assembled by sequentially chemisorbing polyanionic and polycationic materials on
25 miniature (5 x 10" 4 cm 2 ) enzyme electrodes. The sequential chemisorption process
allowed the simultaneous tailoring of their sensitivity, dynamic range, drift and
selectivity. When assembled on tips of 250 (am diameter gold wires coated with a
redox polymer/glucose oxidase sensing layer, they allowed tailoring of the glucose
electrodes for greater than 2 nA/mM sensitivity; 0 to 30 mM dynamic range; drift of
30 <> 5% per 24 hours at 37°C at 15 mM glucose concentration ; < 5% current increment
by the combination of 0. 1 mM ascorbate, 0.2 mM acetaminophen and 0,5 mM urate.
The membranes also retained transition metal ions that otherwise bind to and
WO 01/88524 PCT/US01/15019
12
damage the redox polymer and the enzyme. The electrodes were tested in the jugular
veins and in the intrascapular subcutaneous region of anaesthetized and heparinized
non-diabetic Sprague-Dawley rats, in which rapid changes of glycemia were forced
by intravenous glucose and insulin. After one-point in-vivo calibration of the
5 electrodes, all of 152 data points were clinically accurate when it was assumed that
after insulin injection the glycemia in the subcutaneous fluid lags by 9 minutes
behind that of blood withdrawn from the insulin-injected vein.
EXAMPLES
10 The following are the materials used to make the electrodes: Glucose
oxidase (GOx) (Fluka, Milwaukee, WI, EC 1.1.3.4, 197 units/mg) from Aspergillus
niger\ poly (ethylene glycol) diglycidyl ether (400) (PEGDGE), and polyallylamine
hydrochloride (MW 50,000) (PAm) from Polysciences, Warrington, PA; and
polyacrylic acid sodium salt (PAc) (MW 15,000), l-[3-(dimethylamino)propyl]-3-
1 5 ethylcarbodiimide (EDC), and N-hydroxysuccinimide (NHS) from Aldrich,
Milwaukee, WI. The redox polymer, poly (4-vinylpyridine) partially N-complexed
with [Os(bpy) 2 Cl] +/2+ and quaternized withbromoethylamine (POs-EA), was
prepared as described in Gregg et al., J. Phys. Chem. 95, 5970 (1991), incorporated
herein by reference. The related polymer from which the osmium complex was
20 omitted, PVPEA, was prepared by quaternizing poly(4-vinylpyridine) with
bromoethylamine.
Example 1
Formation of Glucose Electrodes. Miniature gold electrodes were
25 structurally similar to those described in Csoregi et al., Anal. Chem. 66, 3131 (1994).
and U.S. Patent No. 5,593,852, both of which are incorporated herein by reference.
The electrodes were made of polyimide-insulated 0.25 mm gold wire, which was cut
to 5 cm long pieces. At one end, the insulation was stripped from 0.5 cm of the wire
to make the electrical contact. At the other end, a 90 |im deep polyimide-walled
30 recess was formed by electrochemically etching away the gold under galvanostatic
control by an EG&G PARC 273 A potentiostat/galvanostat.
WO 01/88524 PCT/US01/15019
13
The tip of the gold wire at the bottom of the shielded recess was coated with
the transduction (sensing) layer; a nucro-membrane; and a biocompatible layer. The
first and third layers were formed by micropipetting polymer solutions onto the gold
surface under a microscope, using a micromanipulator. The micro-membrane was
5 formed by dip and rinse cycles.
The sensing layer included the redox polymer POs-EA and GOx crosslinked
with PEGDGE. A 20 mg/mL solution of GOx were dissolved in a 0.1 M sodium
bicarbonate aqueous solution. The GOx solution was then mixed at 2: 1 volume ratio
with a 12 mg/mL solution of sodium periodate and the mixture was reacted in the
10 dark at room temperature for 1 hour. 2 jxL of the now periodate-oxidized GOx
solution was mixed with 16 pL of 10 mg/mL POs-EA solution and 1.4 |iL of 2.5
mg/mL PEGDGE solution. 15 droplets of about 5 nL mixed solution were
sequentially micropipetted into the recessed cavity formed by back-etching the gold
in its polyimide insulation. The resulting films were cured at 45°C for 30 minutes.
15 The micro-membrane was formed over the sensing layer. The
polyelectrolyte solutions 20 mM PAc, 20 mM PAm and 20 mM PVPEA (the
concentrations being those of the acidic or the basic functions, not of the
macromolecules) were prepared in 0,1 M NaCl buffered at pH 6 with 0.1 M 2-[N-
morpholino] ethanesulfonic acid (MES). This buffer was used also to prepare 20
20 mM EDC and 50 mM NHS. The sensors were coated by dipping and rinsing cycles,
alternately in PAc and in PAm, to form PAc/PAm bilayers, or in PAc then in
PVPEA to form the P Ac/P VPEA bilayers. The sequence of the resulting sextets
was PAc/PAm/PAc/PAm/PAc/PVPEA; the slashes (/) representing rinses with MES
buffer to remove the excess (unbound) polyelectrolyte. All of the sensors used in
25 vitro and in animals, except for those made for the parametric optimization of the
sensors, had three of the sextet layers (18 layers total).
The biocompatible layer was formed over the micro-membrane. The
biocompatible layer was formed by UV-photocrosslinking tetraacrylated PEO, using
2,2-dimethoxy-2-phenyl-acetophenone as the photoinitiator.
30
WO 01/88524 PCT/US01/15019
14
Example 2
In Vitro Experiments using the Electrode of Example 1. In vitro
experiments were carried out in a stirred, water-jacketed electrochemical cell in 0.15
M NaCl, 0.02 M phosphate buffer solution with pH 7.1. The cell had a saturated
5 Ag/AgCl reference electrode, a platinum counter electrode and the modified 0.25
mm gold wire tip working electrode, as described in Example 1. Unless otherwise
stated the working electrode was poised at 400 mV vs. Ag/AgCl, and the cell was
maintained at 37°C with an isothermal circulator (Fisher Scientific, Pittsburgh, PA).
The potential was controlled by a CHI832 electrochemical detector (CH Instrument,
1 0 Austin, TX) and a PC collected the data.
Figure 1 illustrates the dependence of the sensitivity on the number of
PAc/PAm/PAc/PAm/PAc/PVPEA sextets where the dashed line indicates no micro-
membrane, the squares indicate one sextet, the triangles indicate two sextets, and the
circles indicate three sextets. This demonstrates the expansion of the linear range
1 5 and the corresponding decrease in sensitivity when an increasing number of
countercharged polyelectrolyte layers is applied on the sensing layer. When the
micro-membrane consisted of two sextets (of the sequence
PAc/PAm/PAc/PAm/PAc/PWEA/PA the current
increased linearly with the glucose concentration at least up to 30 mM. As seen in
20 Figure 1, the linear domain in-vitro now extended through the entire physiologically
relevant 2-30 mM range.
Figure 2 illustrates the dependence of the apparent stability on the number of
PAc/PAm/PAc/PAm/PAc/PVPEA sextets where line (a) indicates no micro-
membrane, line (b) indicates two sextets, and line (c) indicates three sextets. Figure
25 2 illustrates the improvement in the stability of the current at 15mM glucose
concentration and at 37°C when the number of the layers was increased. In the
absence of a micro-membrane, 39% of the current was lost in the initial 24 hour
period. Application of two sextets (PAc/PAm/PAc/PAnyPAc/PVPEA/PAc/PAm/
PAc/PAm/PAc/PVPEA) reduced the loss to 9%. When three sextots were applied
30 (PAc/PAm/PAc/PAm/PAc/PWEA/PAc/P
Ac/PAm/ PAc/PVPEA), the 24 hour loss dropped to 5%.
WO 01/88524 PCT/US01/15019
15
Both urate and the ascorbate anions are electrooxidized at potentials positive
of 200 mV (SCE). In the absence of a micro-membrane the electrooxidation current
of anionic interferents is disproportionately high when the redox-polymer backbone
is apolycation. For example, the ascorbate electrooxidation current at 0.1 mM
5 ascorbate concentration is greater than the glucose electrooxidation current at ImM
concentration. The cause of the disproportionate electrooxidation of anionic
interferents is thought to be due to the scaling of their concentration within the redox
polymer with the density of cationic sites. As a result, the permeability of the
membrane to anionic interferents (which is the product of concentration and
10 diffiisivity) is higher than that of neutral molecules like glucose.
Application of the micro-membrane alleviated the disproportionately large
interference by ascorbate and urate, as shown in Figure 3. At 10 mM glucose
concentration, the aggregate increase in current produced by the combination of 0.1
mM ascorbate, 0.2 naM acetaminophen and 0.5 mM urate was less than 5%. The
1 5 sensitivity of the electrodes to glucose was not changed.
Transition metal ions like Zn 2+ and Fe 2+/3+ , which are present in serum at <50
(jM concentration, can reduce the sensitivity of the "wired" enzyme electrode. The
loss is observed even in the presence of 20 mM phosphate, which precipitates most
of these ions at neutral pH. Application of multiple bilayers consisting only of
20 PAc/PAm, did not appear to alleviate this loss. The loss was, however, alleviated by
incorporating PAc/PVPEA bilayers, which had transition-ion complexing pyridine
functions as shown in Figure 4.
The results show that the following sensor characteristics can be
simultaneously provided by a tailored micro-membrane: Linear range, 2 to 30 mM;
25 sensitivity per unit area of 4 to 6 [xAcm" 2 mM~ ! , translating to 2 to 3 nA mM" 1
sensitivity for an electrode having a 5 x 10" 4 cm 2 mass transporting and sensing area;
and < 5% loss in sensitivity in 24 hrs.
Example 3
30 In Vivo Experiments using the Electrode of Example 1. Male Sprague-
Dawley rats, 400-500 g, were pre-anesthetized with halothane (Halocarbon
Laboratories, North Augusta, SC) and anesthetized by intraperitoneal injection (0.5
WO 01/88524 PCTYUS01/15019
16
mL) of a solution made of equal volumes of acepromazine maleate (10 mg/mL),
ketamine (100 mg/mL), and xylazine (20 mg/mL). The animals were shaved about
their necks, abdomens, and between their scapulae, and then secured on a
homeothennic blanket system (Harvard Apparatus, South Natick, MA). A 0.0375-
5 in. -diameter medical grade silicone tube was inserted into the proximal portion of .
their right external jugular vein and secured with 4-0 silk sutures. A dose of 100
units/kg body weight of heparin solution was then administered, followed by an
equal volume of saline to clear the line. A glucose sensor was implanted
subcutaneously between the scapulae, using a 22-gauge Per-Q-Cath introducer
10 (Gesco International, San Antonio, TX). The sensor was taped to the skin to prevent
its movement. A second silicone tubing of -2 cm length was inserted into the
proximal side of the left external jugular vein as a guide, and the second glucose
sensor was inserted inside the guide tube. The tube and the sensor were then secured
with a microvascular clamp, with the sensor protruding beyond the end of the guide
15 tube. A Ag/AgCl surface skin reference electrode was attached to the animal's
abdomen after conductive gel was applied. The sensors and the reference electrode
were then connected to an EG&G PARC Model 400 bipotentiostat, the output of
which was recorded with a Rustrak Ranger data-logger (Rustrak Ranger, East
Greenwich, RT). Data collection started 40-60 minutes after the sensors were poised
20 at +450 mV vs. Ag/AgCl. The reference blood samples were collected from the right
jugular vein and were analyzed using a YSI Model 2300 glucose analyzer (YSI,
Yellow Spring, OH).
In each experiment a 50% glucose solution (300 mg/kg) was administered
intravenously to induce a rapid rise in glucose concentration. A rapid decline in
25 glucose concentration was then induced by an intravenous insulin injection (regular
U-100, 0.5 unit/kg). At the end of the experiment, the rat was euthanized by
intravenous sodium pentobarbital injection, consistent with the recommendations of
the panel on Euthanasia of the American Veterinary Association. The protocols of
. the experiments in vivo were approved by the University of Texas Institutional
30 Animal Use and Care Committee.
The experiment was started when the glycemia of the rats was at a steady
state, the steady glucose concentration being confirmed by withdrawing three blood
WO 01/88524 PCT/US01/15019
17
samples and their analysis with the YSI analyzer. Two minutes after the third
withdrawal a bolus of glucose was injected. The sensors were calibrated in vivo by
independently analyzing a single sample of blood about 2 minutes before the
injection of the bolus of glucose. Figure 5(a) shows the variation in the sensor-
5 measured glycemia after boli of glucose and insulin were sequentially administered
intravenously. Figure 5(b) shows the correlation of the sensor readings and the YSI
results of Figure 5(a). The linear regressions for the eight sensors are summarized in
Table 1, below. The average correlation coefficient (r 2 ) was 0.960 for the jugular
vein sensors and 0.935 for the subcutaneous sensors. The average of the intercept
10 was 0.6mM ± 0.6mM, not differing greatly from the reported -0.79 mM to +0.48
mM range of intercepts of home blood glucose meters used by self-monitoring
diabetic patients.
As seen in Figure 5 and in Table 1 the subcutaneous and the jugular- vein
implanted sensors with the in situ synthesized micro-membranes accurately track the
1 5 YSI-glucose analyzer measured blood glucose concentration when calibrated in vivo
at one point. The clinical validity of glucose assays is often judged by their position
in zones of the Clarke plot. Points in zone A of the Clarke plot represent accurate
assays. Points in zone B represent less accurate assays leading to valid clinical
decision. Points in Zone C reflect assays leading to inappropriate, though not
20 harmful, clinical decisions. Points in zone D reflect assays leading to the missing of
a necessary clinical action (consumption of a glucose-source or insulin-injection)
when such action is required. Points in Zone E reflect assays leading to clinical
action that are the opposite of the required, such as assays indicating the need to
inject insulin when the patient is already hypoglycemic.
25 In absence of correction for the potential transient difference between the
blood and the subcutaneous glucose concentrations during rapid rise or decline
periods the Clarke-type error grid analysis (Figure 6) of the data shows that 95.5% of
the points are in the clinically accurate or acceptable zones A or B (Table 2, below).
The points in zone D (4.5%) resulted of failure to detect hypoglycemia and
30 originated in periods of rapid decline following intravenous administration of
insulin. The fraction of points in zone D was reduced to less than 1% when it was
assumed that following insulin injection, but not after glucose injection, the
WO 01/88524 PCT/US01/15019
18
subcutaneous glucose concentration lags behind that of the insulin-injected vein by 9
minutes. Table 3, below, shows that the fraction of points in zones other than zone
A of the Clarke plot increased when it was assumed that in the period after
intravenous administration of glucose the lag time of the subcutaneous glucose
5 concentration behind that in the glucose-injected vein was greater than 0 to 3
minutes; when no lag or a 3 min lag were assumed all points were in zone A.
The assumptions of a 0 to 3 min lag of the subcutaneous glycemia after
glucose injection and of a 9 min lag after insulin injection not only brought all of the
88 points measured with the subcutaneous electrode into zones A and B of the
1 0 Clarke plot, but also increased the ratio of zone A to zone B points. For the
intravenous sensor-measured glycemia, the assumption of a 3 min lag of the
contralateral venous glycemia behind that in the injected vein, whether after glucose
or insulin injection, brought all points into zones A and B of the Clarke plot (Tables
3 and 4, below). The assumptions of a 9 min lag in the subcutaneous glycemia after
15 insulin injection and of a 3 min lag in the subcutaneous glycemia after glucose
injection, as well as in the contralateral venous glycemia after glucose or insulin
injection, thus brought 163 of the 176 points (92.6%) into zone A; 13 points being in
zone B. Comparison of the values in Table 5, below, with those in Table 1
summarizes the effect of these assumptions on the distribution of points, the slopes,
20 the intercepts, and the percent difference between the YSI and the sensor readings.
The sensors with the in-situ assembled micro-membranes accurately measured the
glycemia in the jugular vein and in the interstitial subcutaneous fluid.
The present invention should not be considered limited to the particular
25 examples described above, but rather should be understood to cover all aspects of
the invention as fairly set out in the attached Claims. Various modifications,
equivalent processes, as well as numerous structures to which the present invention
may be applicable will be readily apparent to those of skill in the art to which the
present invention is directed upon review of the instant specification.
30
WO 01/88524
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PCT/US01/15019
TABLE 1
Results of linear regression analysis of the correlation between the actual
glucose blood glucose concentrations and concentrations measured by the
5 implanted sensors* 0
Slope
Intercept (mg/dL)
r 2
%difference* b
jugular vein
0.83
9.7
0.966
-3.3
jugular vein
0.96
15.0
0.944
11.5
jugular vein
1.02
-4.2
0.957
-2.7
jugular vein
0.94
-0.9
0.972
-9.2
subcutaneous
0.85
13.1
0.943
3.3
subcutaneous
0.90
30.7
0.938
12.7
subcutaneous
0.90
3.5
0.928
-9.0
subcutaneous
0.86
15.7
0.931
0.1
Average
0.91+0.06 10.3±11.1
0.947+0.016
0.4+8.3
*a.. 22 blood samples were withdrawn and independently analyzed in each
experiment.
20 *b. %difference =[£((sensor glucose-blood glucose)/blood glucose)]/n
TABLE 2
Clarke-type error grid analysis of all data, without and with assumption of lag.
25 Zone No lag assumed Lag assumed
Data points
%
Data points
%
A
151
85.8
163
92.6
B
17
9.7
13
7.4
C
0
0.0
0
0.0
D
8
4.5
0
0.0
E
0
0.0
0
0.0
* A 3 min lag was assumed, except for the subcutaneous sensors after insulin
injection, for which a 9 min lag was assumed.
35
WO 01/88524
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20
TABLE 3
Dependence of the distribution of the data points between the zones of the
Clarke plot on the assumed lag after injection of glucose.
5
Time Delay (mins)
0
3
5
7
9
11
Intravenous Sensors
Zone A
32
31 .
30
30
26
24
ZoneB
0
1
2
2
6
7
10
ZoneD
0
0
0
0
0
1
Subcutaneous Sensors
Zone A
32
32
30
30
29
26
ZoneB
0
0
2
2
3
6
15
ZoneD
0
0
0
0
0
0
TABLE 4
Dependence of the distribution of the data points between the zones of the
20
Clarke plot on the assumed lag after injection of insulin.
Time Delay (mins)
0
3
5
7
9
11
25
Intravenous Sensors
Zone A
35
41
40
38
34
33
Zone B
7
3
4
6
10
11
ZoneD
2
0
0
0
0
o
30
Subcutaneous Sensors
Zone A
28
34
36
35
37
34
ZoneB
10
7
6
7
7
10
ZoneD
6
3
2
2
0
0
35
WO 01/88524
PCT/US01/15019
21
TABLES
Results of linear regression analyses of the correlation between the YSI-
measured venous glycemia and the sensor-measured glycemia assuming the
optimal lag times.*
5
Slope
Intercept, mg/dL
r 2
%Difference
jugular vein
0.78
9.8
0.964
12.5
jugular vein
0.99
10
jugular vein
1.03
-11.6
0.981
5.7
jugular vein
0.94
-5.8
0.980
9.6
average
0.94±0.07
2.4J&.6
0.973+0.007
9.4±1.9
subcutaneous
0.85
7.5
0.947
10.7
15
subcutaneous
0.97
-23.8
0.954
16.9
subcutaneous
0.97
3.8
0.975
5.7
subcutaneous
0.88
2.3
0.973
9.0
average
0.92±0.05
-2.5±7.2
0.962±0.012
10.6±3.0
20
average (all data) 0.93 ±0.08
-U±11.6
0.968±0.012
10.0±3.6
* 3 min lag for the contra-lateral venous and the subcutaneous glycemia after
injection of glucose; 3 min lag of the contra-lateral venous glycemia after injection
of insulin; 9 min lag of the subcutaneous glycemia after injection of insulin.
WO 01/88524
WHAT IS CLAIMED IS:
22
PCT/US01/15019
1. A sensor, comprising:
a sensing layer disposed on a substrate;
a multilayer flux-limiting membrane disposed over the sensing layer, the
membrane comprising a first layer disposed on the sensing layer, the first layer
binding to the sensing layer, the membrane further comprising one or more
subsequent layers disposed on and bound to the immediately preceding layer of the
membrane.
2. The sensor of claim 1 , wherein the sensor has an active area of 0. 1
cm 2 or less.
3. The sensor of claim 1, wherein the active area is in a range of 1G" 2
cm 2 to 10" 8 cm 2 .
4. The sensor of claim 1, wherein the active area is in a range of 10" 3
cm 2 to lO^cm 2 .
5. The sensor of claim 1, wherein the sensing layer comprises an
enzyme.
6. The sensor of claim 1, wherein the substrate comprises an electrode
upon which the sensing layer is disposed.
7. The sensor of claim 1, wherein the substrate comprises an optical fiber
upon which the sensing layer is disposed.
8. The sensor of claim 7, wherein the sensing layer is disposed on a tip of
the optical fiber.
WO 01/88524 PCT/US01/15019
23
9. The sensor of claim 1, wherein the sensing layer comprises a non-
leachable redox compound.
1 0. The sensor of claim 9, wherein the non-leachable redox compound
comprises a redox polymer.
1 1 . The sensor of claim 1, wherein the multilayer membrane comprises at
least one polycationic layer.
1 2. The sensor of claim 1 , wherein the multilayer membrane comprises at
least one polyanionic layer.
1 3 . The sensor of claim 1 , wherein the multilayer membrane comprises at
least one polycationic layer and at least one polyanionic layer.
14. The sensor of claim 1, wherein the sensor is adapted for implantation
of at least a portion of the sensor in an animal.
15. The sensor of claim 1, wherein the sensor is adapted for subcutaneous
implantation of at least a portion of the sensor in an animal.
16. The sensor of Claim 1, further comprising a biocompatible layer
disposed over the multilayer membrane.
17. The sensor of claim 1, wherein the membrane comprises at least two
layers alternating between polyanionic material and polycationic material.
1 8. The sensor of claim 1, wherein at least one of the layers of the
multilayer membrane comprises functional groups that can capture transition metal
ions.
WO 01/88524 PCT/US01/15019
24
19. The sensor of claim 1, wherein the membrane is configured and
arranged to reduce flux of an analyte to the sensing layer.
20. The sensor of claim 1 , wherein the membrane is configured and
arranged to reduce flux of at least one interferent to the sensing layer.
21. The sensor of claim 1, wherein the multilayer membrane comprises at
least three layers.
22. The sensor of claim 1, wherein the multilayer membrane comprises at
least six layers.
23. A method of making a sensor, the method comprising steps of:
disposing a sensing layer on a surface; and
forming a membrane over the sensing layer by forming, in an alternating and
binding manner, at least three layers, of which one is a polyanionic layer and two are
polycationic layers or of which two are polyanionic layers and one is a polycationic
layer.
24. The method of claim 23, further comprising disposing a biocompatible
layer over the membrane.
25. A method of making a sensor, the method comprising steps of:
disposing a sensing layer on a substrate;
disposing a first membrane layer on the sensing layer and binding the first
membrane layer to the sensing layer, and
disposing one or more subsequent membrane layers over the first membrane
layer, each subsequent membrane layer binding to the preceding membrane layer.
26. The method of claim 25 , wherein the membrane layers are bound to
the sensing layer or preceding membrane layer by chemisoiption or reactive
adsorption.
WO 01/88524
25
PCT/US01/15019
27. The method of claim 26 , further comprising rinsing after disposing at
least one of the first membrane layer or the one or more subsequent membrane
layers.
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
2/8
PCT/US01/15019
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
3/8
PCT/US01/15019
FIG.3
50 n —
40 -
30 -
0 -I : , ,-J rJ 1 L, L—,-
0 500 1000 1500 2000 2500 3000
Time (sec)
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
4/8
PCT/US01/15019
FIG.4
120 t
Time (sec)
SUBSTITUTE SHEET (RULE 26)
WO 01/88524 PCT/US01/15019
5/8
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
6/8
PCT/US01/15019
FIG.6
YSI-Measured Blood Glucose (mg/dL)
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
7/8
PCT7US01/15019
SUBSTITUTE SHEET (RULE 26)
WO 01/88524
8/8
PCT/US01/15019
CO
o.
CO
o
o
cr
00
CD
SUBSTITUTE SHEET (RULE 26)
INTERNATIONAL SEARCH REPORT
International application No.
PCT/US01/IS019
A. CLASSIFICATION OF SUBJECT MATTER
IPq7) : G01N 27/26
US CL : 2D4./4.I8
According to International Patent Classification (IPC) or to both national classification and IPC
B. FIELDS SEARCHED
Minimum documentation searched (classification system followed by classification symbols)
U.S. : 204/418, 403, 415; 422/82.05, 82.06; 427/2.11, 2.12, 2.13
Documentation searched other than minimum documentation to the extent that such documents are included in the fields
searched
Electronic data base consulted during the international search (name of data base and, where practicable, search terms used)
EAST
search terms: electrode, microelectrode, optical fibre or fiber, membrane, chemisorption, adsorption
DOCUMENTS CONSIDERED TO BE RELEVANT
Category*
Citation of document, with indication, where appropriate, if the relevant passages
Relevant to claim Mo.
US 5,773,270 A (D'ORAZIO et al.) 30 June 1998, the abstract;
Figure 2; column 8, line 36 - col. 9, In. 5; col. 9, 11. 29-49; and col.
10, U. 14-59.
US 5,696,314 A (McCAFFREY et al.) 09 December 1997, the
abstract; Figure 1; col. 4, 11. 37-43; col. 7, In. 65 - col.- 8, In. 27;
col. 10, 11. 31-55; and col. 13, 11. 18-50.
US 6,015,480 A (CRAIG et al.) 18 January 2000, the abstract; and
col. 2, 11. 30-55.
US 5,804,048 A (WONG et al.) 08 September 1998, the abstract;
col. 1, 11. 1-40; and col. 3, 11. 14-45.
1, 19-21
2-6, 11-17, 22-27
1-6 and 9-27
1-6 and 9-27.
1-6 and 9-27
~x| Further documents are listed in the continuation of Box. C. | | See patent family annex.
■bp
"L"
Special caiegorie* of cited documents:
document defining the general state of the art which Is not oonsiderod
to be of particular relevance
earlier document published an ar after the international fi l in g date
document which mar throw donbta on priority olaha(fi) or which h
oiled to establish the poblioation date of another citation cr other
special reason (as specified)
document referring to on oral disclosure, use, exhibition or other
rt moment published prior to the International filing date but later
Pun thn priority data chimed
later document published ifter the intimation al date or priority
date and pol in conflict with the application but oited to understand
the principle or theory underlying thi Invention
document of particular tolerance; the olatmed invention cannot be
considered novel or Caonot be considered to in voire an inventive step
when the document b tahoa alono
document of particular rebmnce; the claimed invention cannot be
ecocide red to Involve an inventive tUp when the document is combined
with one or more other such doc omenta, cuch combination being
abviom to a person skilled in the art
document member of the seme patent family
Date of the actual completion of the international search
16 JULY 2O0I
Name and mailing address of the ISA/ US
Commissioner of Patents and Trademarks
Box PCT
Washington, D.C. 20231
Facsimile No. (703) S05-S2SO
Date of mailing of the international search report
0 4 SEP ZOflt
Authorized officer
ALEX NOGUEROLA CS3QRAH THOMAS
^, , PARALEGAL SPECIALIST
Telephone No. (70S) 308-0661
8
t
Form PCT/ISA/SIO (second sheet) (July 1998)*
INTERNATIONAL SEARCH REPORT
International application No.
PCT/US01/15019
C (Continuation). DOCUMENTS CONSIDERED TO BE RELEVANT
Category*
Citation of document, with indication, where appropriate, of the relevant passages
Relevant to claim No.
US *,943,364 A (KOCH et al) 24 July 1990, the abstract; Figure 1;
and claim 6.
WO 9410553 Al (SINGH) 11 May 1994, the abstract; Figure 2; and
claims 17, 19, 20 -22, 26, 27, and 28.
BP 0352610 A2 (LYBE) 31 January 1990, the abstract; claims 6, 9,
and 1L
US 5,611,900 A (Worden et al) 18 March 1997, the abstract,
Figures l-3b; coL 4, 1L 35-63; coL 6, 1143-50; and coL 7, 11 36-49.
1-5, 7-10, 14-16,
19, 20
1-4, 7, 8, 14-16,
19, 20-22, 25.
1-4, 7, 8, 14-16,
19, 20, 21, 25.
1-6
19, 20
Form PCT/ISA72IO {continuation of second sheet) (July 1998)*
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